1. Background of the Invention
This invention relates to MRI (Magnetic Resonance Imaging) apparatus, and more particularly to orthogonal RF (radio frequency) coils for use in generating high frequency magnetic fields to excite magnetization in the body of an examinee, and detecting NMR (Nuclear Magnetic Resonance) signals generated with mitigation of the magnetization excited.
2. Description of the Related Art
An orthogonal RF coil used in an MRI apparatus has two electrically independent resonator coil loops formed orthogonal to each other. Such a coil is useful in that the two coil loops consume half the high frequency power consumed by a single resonator coil loop when effecting excitation needed for MR imaging, and provide a .sqroot.2times S/N ratio when receiving NMR signals. In order to approximate the above theoretical values, it is important for this orthogonal RF coil to attain a high degree of independence or orthogonality between the two resonator coil loops. The orthogonality of the two resonator coil loops refers to a 90.degree. displacement in vector direction between high frequency (RF) magnetic fields generated by the respective coils in signal transmission, and a 90.degree. displacement in vector direction between NMR signals detected by the respective coils in signal reception.
FIG. 1 shows a first type of known orthogonal RF coil. In FIG. 1, four linear conductors 11-14 are connected at opposite ends thereof to ring-shaped conductors 15 and 16 through capacitors 31-34 and 35-38. The ring-shaped conductors 15 and 16 include capacitors 21-24 and 25-28 arranged in series, respectively. The linear conductors 11 and 13, ring-shaped conductors 15 and 16, and capacitors 31, 33, 35 and 37 form one resonator coil loop, while the linear conductors 12 and 14, ring-shaped conductors 15 and 16, and capacitors 32, 34, 36 and 38 form the other resonator coil loop. Power is supplied to points B and C from a 90.degree. phase power divider 51. Capacitors 39 and 40 are provided to effect impedance matching.
In this example, the capacitors 21-28 have a relatively large capacity, with mainly capacitors 31-38 acting to adjust resonance frequencies to obtain a steady characteristic (i.e. maintenance of the orthogonality noted above). Each of the capacitors in the groups of four capacitors 21-24, 25-28, 31-34 and 35-38 have substantially the same capacity.
Resonance currents as shown in FIG. 2A flow through this type of orthogonal RF coil. In FIG. 2A, resonance currents shown in solid lines and those shown in dotted lines flow independently of each other. The current flowing through point C and the current flowing through point A are in phase inversion. Similarly, the current flowing through point B and the current flowing through point D are in phase inversion. These independent currents generate RF magnetic fields orthogonal to one another. Consequently, it is useful in MR imaging to establish a resonance mode with such current flows at NMR resonance frequencies (imaging mode).
On the other hand, there exists also a resonance mode (differential mode) with resonance currents flowing as shown in FIG. 2B. In this mode, the currents flow commonly through linear conductors 11-14, with no independence thereamong. The current flowing through point C and the current flowing through point A are in phase, and so are the current flowing through point B and the current flowing through point D. A composition of RF magnetic fields centrally of the RF coil is zero, and so this mode generally cannot be used as an imaging mode.
However, the known orthogonal RF coil has a drawback of providing a poor orthogonality of RF magnetic fields. That is, the frequency establishing the imaging mode and the frequency establishing the differential mode often become close to each other because of the balance of capacitance and inductance in the RF coil or the like. In such a case, the resonance currents in the differential mode are combined with the resonance currents in the image mode, thereby degrading the orthogonality of RF magnetic fields in the imaging mode.
Known cylindrical RF coils as described above include what is known as a birdcage coil and a distributed cosine coil. Each of these cylindrical RF coils has numerous linear conductors extending parallel to a cylinder axis. In a resonant state, each linear conductor receives a cos 8 current corresponding to a spatial position 8 thereof, thereby generating high frequency magnetic fields of excellent uniformity (see Japanese Patent Publication 3-50541 for the birdcage coil, and Bolinger et al., J. Magnetic Resonance, 81, 162, 1988 for the distributed cosine coil).
FIG. 3 shows an example of birdcage coils having eight linear conductors. In FIG. 3, the eight linear conductors 111-118 are connected at opposite ends thereof to two ring-shaped conductors 181 and 182. In a resonant state, highly uniform magnetic fields are generated when the eight linear conductors 111-118 form a current distribution as shown in FIG. 4.
As shown in FIG. 5, a distributed cosine coil has linear conductors 211-230 each connected at one end (upper end) thereof to a ring-shaped conductor 281. The other ends (lower ends) of linear conductors 211-230 are joined to form groups (groups of five each in the drawing) to be connected to a ring-shaped conductor 282 through capacitors 251-254. In the drawing, the dotted lines represent an orthogonal type construction (with solid line portions and linear portions generating magnetic fields at 90.degree. to one another).
However, the conventional RF coils have the problems that the uniformity of magnetic fields is impaired as a result of frequency adjustment, and that adjustment to keep the uniformity of magnetic fields is difficult.
Specifically, for use in an MRI apparatus, it is necessary to bring the resonance frequency of the RF coil into agreement with NMR resonance frequency, and tuning capacitors must be adjusted. In the example shown in FIG. 3, the capacitors 131-138 are selected while maintaining the capacity thereof, to adjust the resonance frequency. To maintain a high degree of uniformity of magnetic fields, the capacitors corresponding in number to the linear conductors must be varied while maintaining the capacity thereof equal as noted above. Such an adjusting operation is extremely difficult. If frequency adjustment is effected by varying only certain capacitors to simplify the operation, the distribution of currents flowing through the linear conductors will deviate from what is shown in FIG. 4 and impair the uniformity of magnetic fields.
With the distributed cosine coil shown in FIG. 5, the current distribution is determined only by inductance of each linear conductor. It is therefore difficult to obtain an ideal current distribution, and the current distribution cannot be adjusted once a shape has been determined. That is, the currents flowing through the linear conductors 211, 212 and 213, for example, are determined by line lengths up to points 272, 273 and 274, respectively. It is thus impossible to effect adjustment for enhancing the uniformity of magnetic fields. Further, the linear conductors 211-215 and ring-shaped conductor 281 form a direct current type closed loop. This closed loop is guided by a pulse-like application of gradient magnetic fields during MR imaging, to generate over-currents detrimental to reconstructed images.